The signal processing described herein is primarily described in the context of emission tomography medical imaging. The processing, however, can be utilized in connection with many other types of scintillation systems (e.g., well logging systems). Therefore, the discussion regarding medical imaging describes, by way of example, one of the many modalities in which the signal processing may be implemented.
One known type of emission tomography medical imaging is generally known as positron emission tomography (PET). PET scanners are utilized to generate images of, for example, portions of a patient's body. Positron annihilation events are utilized in generating such images. Positrons (a positron is the antiparticle of the electron) are emitted by radionuclides prepared using a cyclotron or other device. The radionuclides are employed as radioactive tracers called "radiopharmaceuticals" by incorporating them into substances, such as glucose or carbon dioxide.
The radiopharmaceuticals are injected into the patient and become involved in such processes as blood flow, fatty acids, glucose metabolism, and synthesis. Positrons are emitted as the radionuclides decay. The positrons travel a very short distance before encountering an electron, and when that occurs, the position and electron annihilate emitting two photons directed in nearly opposite directions.
In some known PET scanners, two detector heads located one hundred and eighty degrees apart rotate around a patient. Each detector head includes crystals, referred to as scintillators, to convert the energy of each 511 keV photon into a flash of light that is sensed by a photomultiplier tube (PMT). Coincidence detection circuits connect to the detectors and record only those photons that are detected simultaneously by the detector. The number of such simultaneous events indicates the number of positron annihilations that occurred along a line joining the two opposing detectors.
During a scan, hundreds of millions of events are detected and recorded to indicate the number of annihilation events along lines joining pairs of detectors in the ring. The collected data is used to reconstruct an image. Further details regarding PET scanners are set forth in U.S. Pat. Nos. 5,378,893, 5,272,343, and 5,241,181, all of which are assigned to the present assignee.
Generally, by maximizing the count rate (i.e., the number of detected events), the performance of the imaging system is enhanced. The anode signals from the PMTs must be processed to provide a good measurement of the integrated signal strength. Two known methods typically are utilized for processing the anode signals. In one method, the signal pulse is shaped by a filter (RC-CR, Gaussian, etc.), and the peak value of the shaped pulse is then digitized. The pulse shaping method provides the advantage of low noise and produces an output without the necessity of a trigger circuit. However, the shaped pulse is wide and the pile up of the filtered/shaped pulses limits the usefulness of the circuit at high count rates.
The other known method is referred to as a switching integrator method. Specifically, a timing pickoff circuit is used to detect the leading edge of the signal pulse. The output of the timing pickoff circuit then initiates operation of an integrator (either analog or digital) which integrates the signal pulse. The switching integrator method does not require that the pulse be broadened (filtered). Therefore, this method is less affected by pileup at high count rates. However, the circuit does not produce an output unless a trigger was detected and the circuit is dead during the integration, which adversely impacts the count rate.
To reduce the effect of the dead time of the integrator on the overall system dead time, multiple integrators can be used on each signal. If a pulse is detected while one integrator is processing an event, the anode signal is switched to another integrator which is not in use. The integration of the first pulse can be stopped when a second pulse is detected and output of the integrator corrected for the shorter integration time. The output value from the first integrator can be used to correct the output value of the second integrator for the contribution from the pileup of the first pulse with the second pulse.
Operating a nuclear camera in a 511 keV coincidence detection mode requires that the detector heads function at as high a count rate as possible with minimum dead time and pileup. It would therefore be desirable to combine the advantages of the pulse shaping method (i.e., continuous output and minimum dead time) with the advantages of the integrator method (i.e, minimum pileup and the ability to easily correct pileup between pulses).